Medical doppler ultrasound system for locating and tracking blood flow

ABSTRACT

Systems and methods for processing echo signals in a Doppler ultrasound system from a region of interest. An ultrasound beam is electronically steered to deliver ultrasound to and receive echo signals from a plurality of sample locations in the region of interest. The echo signals for each sample location are processed to extract Doppler shift signals and Doppler shift data representing the Doppler shift signals are generated. The Doppler shift data accumulated for the sample locations can be used to detect the presence of blood flow in the region of interest, and identify the location in the region of interest at which the presence of blood flow is detected. The blood flow can be tracked by updating the location of the detected blood flow in the region of interest. The blood flow can be further monitored by combining the locating and tracking functionality with an m-mode ultrasound image.

CROSS-REFERENCE TO RELATED APPLICATION

This application is a continuation of U.S. patent application Ser. No.11/152,666, filed Jun. 13, 2005. This application is incorporated byreference herein in its entirety and for all purposes.

STATEMENT AS TO GOVERNMENT RIGHTS

The disclosed invention was made with support from the United StatesGovernment, which has certain rights in the invention pursuant to GrantNo. 1R43NS046055-01 awarded by the National Institutes of Health.

TECHNICAL FIELD

This invention relates generally to medical monitoring and diagnosticdevices, and more particularly, to a medical Doppler ultrasoundapparatus and method for automatically locating and tracking blood flowto enable monitoring of the blood flow in a patient.

BACKGROUND

The devastation that stroke inflicts is easily found in the medicalliterature: stroke strikes someone in the United States every 45 secondsand kills someone every three minutes. Each year about 700,000 peoplesuffer a stroke. About 500,000 of these are first attacks. Of allstrokes, 88 percent are ischemic, nine percent are intracerebralhemorrhage, and three percent are subarachnoid hemorrhage. TheFramingham Heart Study showed that, six months after their strokes, 50percent of ischemic stroke survivors studied over age 65 had somehemiparesis, 30 percent needed help walking, 26 percent were dependentfor daily living, and 19 percent had aphasia. Twenty-six percent wereinstitutionalized. Eight to twelve percent of ischemic strokes are fatalwithin 30 days. As the mean age of the population increases, theincidence of stroke is projected to increase. Stroke clearly compromisesquality of life for its victims and for society at large. In economicterms, the estimated total cost of stroke in the United States in 2004is $53.6 billion. Direct costs include $26.5 billion for hospital andnursing home stays, $2.7 billion for physicians and other professionals,$1.1 billion for drugs and other medical durables, and $2.7 billion forhome health care. Indirect costs are estimated at $6.1 billion in lostproductivity due to morbidity and $14.5 billion in lost productivity dueto mortality. For Americans age 40 and older, 1995 data showed theaverage in-hospital and physician costs were $11,010 for a stroke and$4,940 for trans-ischemic attack (“TIA”).

Although transcranial Doppler (“TCD”) ultrasound has been available formany years as a standard diagnostic modality, and has been shown to haveutility in the hands of a skilled user for assessing and monitoring thebasal cerebral arteries in the early stroke patient, it has not beenwidely used in this capacity. This is in part because the demand forthis sort of assessment is a recent phenomenon accompanying the adventof thrombolytic therapy, and in part because TCD is difficult at bestfor the emergency medicine physician or nurse to perform and interpret.In elderly populations, where the technology is greatly needed,frustrating time is consumed in finding signals in many patients.Locating the signal is difficult because of the need to laterallyexplore each depth one step at a time with single-gate TCD equipment,until flow signals are acquired. Searching for flow can be a tedious andperplexing task. Once a flow signal is acquired, it is confirmed byconsidering the associated depth, the approximate aim of the probe, andtracing the signal to adjacent vessels as depth is varied to verify theuser is indeed on the appropriate vessel. Successful utilization ofsingle gate analog transcranial Doppler is not generally availableoutside vascular laboratory personnel, and even these ultrasound expertshave very mixed reactions to TCD because of the above mentioneddifficulties. In summary, these factors combine so that there is a lackof utilization of TCD capabilities in triage and monitoring, which iswitnessed by the relatively minor presence of TCD in emergencydepartments across the United States.

The benefits of rapid and easy to perform assessment of cerebralhemodynamics in the early presentation of a patient with suspectedstroke are tremendous, both for the admitting physician and the patientfacing a potentially debilitating or fatal stroke. For the physician,there is the early recognition of pathology from hemodynamicobservations, such as ischemic blockage, its location, and theaccompanying option of thrombolytic therapy. Studies have shown that anuntreated occlusion of the middle cerebral artery presents a very poorprognosis for the patient. Hence, knowing about it as soon as possiblemaximizes the potential for positive intervention. For example, afterinitiating thrombolytic therapy, monitoring can determine the point intime at which thrombolytic therapy re-establishes perfusion, presentingthe possibility of termination of successful thrombolysis in order tominimize risk of bleeding associated with thrombolytic drugs. For thepatient, “time is brain.” The success of aggressive therapy such asthrombolysis is partly dependent on its application in the first threehours after stroke onset. These benefits will be appreciated in theemergency department and by the patient to the degree that theassessment of cerebral hemodynamics is rapid and easy to perform.

Recently, a digital Doppler platform has been developed by SpencerTechnologies in Seattle, Wash. in which up to 33 sample gates can besimultaneously processed into a “color” m-mode image. The color in them-mode image is a function of Doppler signature power and detectedvelocity, in that increases in backscattered power cause the colors, redor blue, to become more intense. The digital Doppler platform isreferred to as Spencer Technologies' Power M-mode Doppler (“PMD”).Showing power in this fashion conveys to the user when the Doppler beamis well aimed—that is, intensity of color increases with volume ofmoving blood in the Doppler sample volume and this indicates when thebeam is centered on the blood flow. Thus, the color m-mode display of anultrasound system having PMD capability provides medical professionalswho do not have expertise in ultrasound with a mechanism for easylocation (by the operator) of the middle cerebral circulation. A moredetailed description of PMD ultrasound systems can be found in U.S. Pat.No. 6,196,972 to Moehring, issued Mar. 6, 2001 and assigned to SpencerTechnologies.

Displaying color as a function of signal power at multiple depths offersadvantages for an examiner to locate a temporal bone window, or an“acoustic window,” without limiting interrogation to a single selecteddepth. When employed in assessment of patients at various vascularlaboratories, the sonographers reported that PMD TCD was easier to usesince it was no longer necessary to seek a window by changing depth.Also they found it unnecessary to listen for a Doppler sound. In fact,an ultrasound system having PMD capability enabled them to first findthe optimal temporal window on the PMD display and then to adjust thegate depth for audible spectral display using the PMD depth scale. ThePMD system allowed the examiner to find the temporal window withoutaudible Doppler sounds and therefore avoid crashing sounds of probeapplication to the head. During intraoperative monitoring, thesonographers maintained their position on the window using the colorsignals present in the PMD image as a guide. When the flow signals wereaccidentally lost during surgical monitoring, these could be recoveredby repositioning the transducer using the PMD display as the onlyfeedback. Therefore, the significant problem of single gate TCD, thatis, frustration in locating difficult windows, was reduced.

Although PMD provides an easy method for aiming the Doppler probe, thetask of locating acoustic windows and underlying blood flow is stillleft to the operator when using the PMD system. Thus, although thedevelopment of the PMD platform mitigated the problems related toneeding highly skilled operators to operate single-gate TCD equipment,the problems are nevertheless still present to some degree.

SUMMARY

One aspect of the invention provides a system and method for processingecho signals in a Doppler ultrasound system from a region of interest.An ultrasound beam is electronically steered to deliver ultrasound toand receive echo signals from the region of interest. The regionincludes a plurality of sample locations. Each location has anassociated beam axis and at least two different planes are defined inwhich two or more of the beam axes lie. The echo signals for each samplelocation are processed to extract Doppler shift signals and Dopplershift data representing the Doppler shift signals are generated.

In another aspect of the invention, a system and method for locatingblood flow in a region of interest using a Doppler ultrasound systemhaving a multi-element transducer is provided. An ultrasound beam issteered to deliver ultrasound to and receive echo signals from aplurality of sample locations. The plurality of sample locations arespatially arranged across the region of interest in two lateraldimensions relative to a reference beam axis extending from themulti-element transducer. For each sample location, echo signals areprocessed to extract Doppler shift signals and generate Doppler shiftdata, which is accumulated for the plurality of sample locations. Basedon the Doppler shift data accumulated for the plurality of samplelocations the presence of blood flow in the region of interest isdetermined, and the location in the region of interest at which thepresence of blood flow is detected is identified.

In another aspect of the invention, a system and method for monitoringblood flow in a region of interest using a Doppler ultrasound systemhaving a multi-element transducer is provided. Blood flow in the regionof interest is located and a vector identifying the location of bloodflow is defined. Ultrasound is delivered to the location of the bloodflow and first echo signals are received from the location of the bloodflow in accordance with the vector. The first echo signals from thelocation of the blood flow are processed to determine Doppler signalstrength and a blood flow velocity toward the probe. The vectoridentifying the location of the blood flow in the region of interest isupdated and ultrasound is delivered to the location of the blood flow inaccordance with the updated vector. Second echo signals are receivedfrom the location of the blood flow in accordance with the updatedvector and are processed to determine Doppler signal strength and ablood flow velocity toward the probe.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is a diagram of a virtual surface which is transected by a bloodvessel and over which an ultrasound beam is scanned. FIG. 1B is a planview of the virtual surface of FIG. 1A. FIG. 1C is a diagram of a powerm-mode image of blood flow of the blood vessel transected by the virtualsurface of FIG. 1A.

FIG. 2 is a diagram of a virtual surface constructed in accordance withan embodiment of the present invention.

FIGS. 3A and 3B are diagrams showing ultrasound echoes from ultrasounddelivered from an ultrasound probe positioned at two locations on theskull of a patient.

FIG. 4A is an isometric drawing of an ultrasound probe and FIG. 4B is aplan drawing of a multi-element transducer according to an embodiment ofthe invention included in the probe of FIG. 4A.

FIGS. 5A-5D are acoustic intensity maps of steered ultrasound beamsdelivered by the probe and multi-element transducer of FIGS. 4A and 4B.

FIG. 6 is a functional block diagram of a Doppler ultrasound systemaccording to an embodiment of the invention.

FIG. 7 is a functional block diagram of a digital signal processing(“DSP”) platform according to an embodiment of the invention.

FIG. 8A is a functional block diagram of a transmit circuit of the DSPplatform of FIG. 7 according to an embodiment of the invention. FIG. 8Bis a functional block diagram of a receive circuit of the DSP platformof FIG. 7 according to an embodiment of the invention.

FIG. 9 is a flow diagram of a process for Doppler shift signalprocessing according to an embodiment of the invention.

FIG. 10 is a flow diagram of a process for locating an acoustic windowaccording to an embodiment of the invention.

FIG. 11 is a flow diagram of a process for locating blood flow ofinterest according to an embodiment of the invention.

FIG. 12 is a diagram showing relationships between various ultrasoundtransducer parameters.

FIGS. 13A-13C are plan drawings of multi-element transducers accordingto embodiments of the invention.

FIGS. 14A-14D are diagrams of visual feedback provided to an operator inaccordance with an embodiment of the invention during a process oflocating blood flow in a region of interest and evaluating the bloodflow.

DETAILED DESCRIPTION

When monitoring motion with Doppler ultrasound, such as monitoring bloodflow or tissue motion, it is desirable to obtain information thatprovides high time resolution of acquired signals and has minimal or notime gaps when observing the motion of interest. This generally occurswhen the ultrasound beam is aimed in the same direction as the bloodflow or tissue motion. As a result, non-scanning ultrasound devices aretypically used in these applications where the ultrasound is not used toscan other regions, but is continually aimed at the region of interestduring monitoring. Examples of non-scanning ultrasound devices are powerM-mode Doppler ultrasound devices, which are described in U.S. Pat. Nos.6,196,972 and 6,616,611 to Moehring, incorporated herein by reference.

Although observing blood flow or tissue motion can be accomplished witha non-scanning ultrasound device, as previously discussed, a first stepof actually locating the target typically requires considerable skill inoperating the equipment. As a result, the time required to initiallylocate the target for monitoring can vary signficantly based the abilityof the operator. For example, in a TCD application, an understanding ofwhere to search for an acoustic window in the skull, and the appropriateaiming of the probe during this search process to discover underlyingblood flow is highly dependent on the skill and experience of theoperator since the operator has little feedback indicating the directionand position of the ultrasound beam relative to the blood flow.

Embodiments of the present invention are directed to a Dopplerultrasound system and a method for locating and tracking blood flow ortissue motion of interest. Included are various embodiments that can beused in TCD applications as a tool for rapid detection of acousticwindows in the skull and underlying blood flow that feeds the cerebralcirculation. Generally, however, these embodiments, as well as others,can be used to locate blood flow or tissue motion of interest, and thenobserve or monitor blood flow or tissue motion continuously. Otherembodiments further provide feedback that can be easily understood by anoperator to assist in locating and tracking blood flow or tissue motionof interest. It is foreseen that various embodiments of the presentinvention can be utilized in various applications and environments toquickly locate and track blood flow of interest, such as, monitoringblood flow in basal arteries in the brain, including the middle cerebralartery, the anterior cerebral artery, and the posterior cerebral artery,the carotid siphon and/or the ophthalmic artery as seen through theorbit, the vertebral arteries and the basilar artery as seen through theforamen magnum. For example, using an ultrasound system according to oneembodiment of the invention, it is possible for an anesthesiatechnologist to apply a probe preoperatively and the anesthesiologist,during the surgical procedures, to utilize the information obtained bythe system in managing cerebral perfusion and advising the surgeon ofdangerous embolic activity. Another example application is in thevascular lab setting, where diagnostic tests such as microembolismmonitoring, evaluation of collateral channels, head-turn syncope, bubbletest for patent foramen ovale, evaluation of stenosis in any majorvessel supplying the brain with blood, and vasospasm mapping may befacilitated using an ultrasound system according to an embodiment of theinvention. In other embodiments, tissue motion of interest can bequickly located and tracked, such as monitoring motion of the tympanicmembrane, and assessing/monitoring brain tissue motion.

Urgent clinical settings, such as the emergency department, the hospitalpatient rooms, the operating rooms, and the vascular laboratory, mayalso benefit from a Doppler ultrasound system according to an embodimentof the invention. For example, an acute stroke patient can be followedfrom the emergency department to the hospital room without interruptingthe monitoring of cerebral arterial blood flow. In the hospitalenvironment nurses and other paramedical personnel can be taught toapply the probe and interpret the information along with othermonitoring parameters. During cardiopulmonary bypass, carotidendarterectomy, and orthopedic surgery, the ability to locate and tracka signal from the middle cerebral artery will eliminate the need for aspecial technologist whose cost has been an impediment to use of TCDhemodynamic and embolic information.

Certain details are set forth below to provide a sufficientunderstanding of the invention. However, it will be clear to one skilledin the art that the invention may be practiced without these particulardetails. Moreover, the particular embodiments of the present inventiondescribed herein are provided by way of example and should not be usedto limit the scope of the invention to these particular embodiments. Inother instances, well-known circuits, control signals, timing protocols,and software operations have not been shown in detail in order to avoidunnecessarily obscuring the invention. Embodiments of the inventiondescribed below are directed to application in monitoring cerebral bloodflow. However, some or all of the aspects of embodiments describedherein can be utilized for monitoring blood flow, generally, as well asmonitoring tissue motion of interest.

Several embodiments are described below to illustrate various aspects ofthe present invention. For example, images of compound-mode Dopplervirtual surfaces that can be used to obtain visual feedback informationfor locating blood flow or tissue motion of interest are describedbelow. One compound-mode Doppler virtual surface described isconstructed from a plurality of sample locations distributed over aregion of interest. The sample locations are interrogated byelectronically steering ultrasound in the direction of the differentsample locations. A multi-element transducer and a Doppler ultrasoundsystem having multi-channel Doppler digital signal processing providingelectronic ultrasound beam steering is also described below. Followingthe description of the ultrasound system, an algorithm for locatingblood flow or tissue motion of interest is described, along with adiscussion of a minimum variance estimation technique that can beutilized in a locating algorithm. Additionally, an algorithm forlocating an acoustic window through which cerebral blood flow can bemonitored is also described and the locating algorithm is described forthe specific application of locating cerebral blood flow. An algorithmfor tracking the located blood flow is described in combination with thelocating algorithm. However, the tracking aspect does not need to becombined with the locating aspect. A discussion of various embodimentsof multi-element transducers follows the discussion of the locatingalgorithm, which is followed by a description of an embodiment thatcombines various aspects of the invention to provide locating, tracking,and monitoring of blood flow.

In one embodiment of the invention, information related to a spatialrelationship between a PMD image, such as that provided in the Dopplerultrasound device described in the aforementioned U.S. Pat. Nos.6,196,972 and 6,616,611, and an image of a compound-mode Doppler“virtual” surface is provided to an operator. Compound-mode is an olderterm in the development history of medical ultrasound and refers to“compound scan” imaging, which has been used to observe tissue at afixed depth from the probe. The surface is virtual in that the surfacedoes not physically exist, but represents a region to which ultrasoundis delivered and from which Doppler information is obtained.

An example of the spatial relationship between PMD, a style of m-modeimaging, and an image of a compound-mode Doppler virtual surface in thecontext of observing blood flow in the middle cerebral artery (“MCA”)120 is illustrated in FIGS. 1A-1C. In FIG. 1A, an ultrasound beam (notshown) emanating from an ultrasound probe 110 has a central axis 112.Various locations along the axis 112 can be sampled for blood flowinformation by evaluating the echoes for motion at points distributedall along the axis 112. As shown in FIG. 1A, a constant-depth sphericalshell at 50 mm depth having a fixed arc length spanning about 20 mmintersects the MCA 120. The shell represents a compound-mode Dopplervirtual surface 116 that is interrogated by ultrasound to obtaininformation for a region of interest. In one embodiment on theinvention, the virtual surface 116 is scanned for Doppler (blood flow)signals and a circular image 150 representative of the virtual surface116 presented to the user, as shown in FIG. 1B. When blood flow from theMCA 120 intersects the virtual surface 116, a spot representing a bloodflow signal 152 appears on the circular image 150. The blood flow signal152 has a color (color not shown in FIG. 1B) according to the directionof motion and has an intensity proportional to the strength of theDoppler signals. For example, a red color can be used to represent bloodflow toward the probe 110 and a blue color can be used to representblood flow away from the probe 110. A PMD image 140, shown in FIG. 1C,is constructed for the blood flow signal 152 and illustrates blood flowat a 50 mm depth over time in a PMD format. As will be explained in moredetail below, the use of a compound-mode Doppler virtual surface 116 canalso be used to quickly locate a good acoustic window through bone andlocate blood flow.

In the present embodiment, a compound-mode virtual surface 116 having a20 mm diameter is provided at a 50 mm depth. To provide coverage over a20 mm diameter, beam steering is employed to detect blood flow at pointsacross the virtual surface 116. A relatively narrow ultrasound beam issteered to various “look directions” 210 across the virtual surface 116,which are shown in FIG. 2. Generally, each of the look directionscorresponds to a sample location in the region of interest. Informationobtained from a series of transmit bursts at each of a series of thelook directions 210 is used to detect blood flow passing through thevirtual surface 116. The series of look directions 210 essentially“tile” the virtual surface 116 so that an array of points in the 20 mmvirtual surface 116 are interrogated with ultrasound. The data which isacquired at a given look direction 210 is constructed from a series oftransmit bursts, and can be used to calculate the angular spectrum ofthe Doppler signal, as will be described in more detail below. Theangular spectrum of the Doppler signal in the local neighborhood of agiven look direction 210 can be used to process greater angularresolution for flow signal location, thereby adding detail to the imageof the virtual surface 116 greater than that illustrated in FIG. 2.

The data from the look directions 210 are processed and combined in amontage that displays the data concurrently. While the image isdisplayed, new data can be acquired and processed to update the image.Data acquisition sufficient to construct an image that providesreal-time feedback to an operator holding the probe is desirable.Generally, updating an image of the virtual surface 116 at ten framesper second, which corresponds to one frame every 100 ms, should provideacceptable feedback and should further provide sufficient time to obtaindata from enough different look directions 210 to adequately survey thevirtual surface 116.

As shown in FIG. 1B, the surface 116 is a circular spherical shell at aconstant radius from the probe 110. However, the surface 116 can havedifferent surface features and shapes without departing from the scopeof the present invention. For example, in one embodiment, the surface116 can be rectangular shaped rather than circular. The surface 116 canbe planar (variable radius from the probe 110) rather than a sphericalshell (constant radius from the probe 110). The surface 116 can betilted with one edge at a different distance from the probe 110 than theopposite edge. The surface 116 may be curved in one-dimension and flatin the second dimension.

Additionally, the look directions, which generally define samplelocations of the surface 116, have been shown having a certain dimensionand spatial arrangement on the surface 116. Both of thesecharacteristics can be modified without departing from the scope of thepresent invention. It will be appreciated, however, that it is desirablefor the sample points to have sufficient density to survey the surface116 for blood flow in the region of interest. Generally, the samplelocations of the surface 116 have an associate beam axis that generallydefines a direction of the sample location. As a result of distributingthe sample locations over a region having lateral extent intwo-dimensions relative to the probe, there are at least two differentplanes that are defined which have two or more beam axes lying in therespective plane. In contrast, B-mode ultrasound is accomplished withmore than one sample point (indeed, many) per beam steering direction orsource position. The depth direction of B-mode produces one dimension ofthe B-mode image at no additional time expense than listeningsufficiently long to echoes from each outgoing pulse. The scanning orimaging forming the surface 116 utilizes two degrees of lateral orangular freedom and generally requires more ultrasound pulses to scanthe surface 116.

The value assigned to a sample point on the virtual surface 116 is afunction of the ultrasound echoes sampled from the ultrasound beamdirected to that particular location. As will be explained in moredetail below, the value is derived from the extracted Doppler shiftsignals along the beam, and may be signal amplitude (or power) at thedepth of the surface 116, the time rate of change of phase of the signal(velocity) at the depth of the surface 116, a frequency domain propertysuch as integrated power associated with velocities above a thresholdand at the depth of the surface 116, or a maximum velocity detectedalong the beam direction, or a combination of some or all of the above.The value assigned may be a scalar or a vector. The value assigned to asample point may subsequently be used in forming an image to bedisplayed to an operator, or subsequently be used for automatic locationor monitoring or tracking of blood flow, as will be described below. Ifused to form a graphical image for the user, the value may do so indifferent ways, for example, a scalar may be depicted simply by imageintensity, while a vector may be depicted by image intensity and color.

In the present embodiment, the virtual surface 116 is constructed from aplurality of look directions 210 in part due to constraints in usingultrasound in a TCD application. That is, in the present example ofdetecting blood flow in MCA, which is typically at a 50 mm depth whenthe probe is on the side of the head at the temporal bone, various beamcharacteristics of the ultrasound beam are a function of the dimensionof the probe at the surface of the skull. A narrow beam diameter at 50mm depth results from constraints of using ultrasound in TCD anddictates the number of look directions to “tile” the virtual surface116. For example, one constraint is that of maintaining a thermal indexcranial (“TIC”) at a level below that of concern for increasedtemperature in the temporal bone. The constraint predicates using lowamplitude ultrasound combined with a wider exposure area (transducerarea) to accomplish a useful acoustic intensity at the depth ofinterest. As a result, a beam having a relatively narrow beam profile isgenerated. Additionally, using an ultrasound beam having a diameter atthe depth of interest that is similar to or smaller than the crosssection of blood flow to be monitored is desirable. A 13 mm diameterprobe transmitting at 2 MHz ultrasound has a transverse dimension at 50mm depth that is similar to the diameter of the MCA, which is typicallyabout 3 mm in diameter. However, the resulting narrow beam isinappropriate for assessing the entire virtual surface 116 with a smallnumber of transmit bursts and is generally limited to being aimed inessentially one direction at a given time. Thus, a series of transmitbursts at a plurality of look directions 210 are used to detect bloodflow across the virtual surface 116.

The ultrasound data acquired at each of the look directions 210 isinterpreted and categorized into three categories for display: (1) nosignal is present because of lack of ultrasound transmission throughbone, (2) a signal present because ultrasound is successfullytransmitted through bone, but there is no blood flow detected, and (3)blood flow is detected. The first category corresponds to a case wherethe look direction 210 is not through an acoustic window. The secondcategory corresponds to the case where the look direction 210 is throughan acoustic window, but the presence of tissue is detected in, theparticular look direction 210 and not blood flow. Each of the threecategories can be displayed differently in an image of the virtualsurface 116. As shown in FIG. 2, the first category is shown as a blackspot on the image, for example, look direction 220. The second categoryis shown as a gray spot on the image, as for look direction 230. Thethird category is shown as a spot on the virtual surface 116 having acolor based on the direction of blood flow detected, as for lookdirections 240 and 250 (color not shown in FIG. 2).

In identifying which look directions 210 are aimed through an acousticwindow or not, the ultrasound reflections are analyzed. FIGS. 3A and 3Bshow examples of ultrasound reflections corresponding to between 40 and60 mm depth from the probe. FIG. 3A shows reflected ultrasound signalswhen the probe is placed on the forehead of a human subject, which is anexample of positioning the probe over a region that is not an acousticwindow. As shown in FIG. 3A, there is no ultrasound transmission throughthe bone and a poor reflected signal is detected at the depth ofinterest. In contrast, FIG. 3B shows reflected ultrasound signals forthe case when the probe is placed on the temporal bone, which typicallycorresponds to a location over or near an acoustic window. As shown inFIG. 3B, strong ultrasound reflections are observed, with the signalshown representing the convolution of the ultrasound transmit burst withthe tissue backscatter function from a region including brain tissue.

In categorizing the reflected ultrasound signals, as previouslydiscussed, the two signals can be differentiated from each other using avariety of techniques, including energy thresholding. For example, ifthe signal level at a desired depth falls below a threshold, the datafrom the look direction 210 can be categorized in the first category.The corresponding look direction 210 can be colored in the image of thevirtual surface 116 to indicate that there is no transmission throughbone, which is black in the present example. However, if the signallevel at a desired depth is above a threshold, then the data from thelook direction 210 can be categorized in the second category and can becolored in the image to indicate transmission through bone, which isgray in the present example. If the signal level is above a thresholdand there is Doppler signal energy which is also above a threshold, thedata from the look direction 210 can be categorized in the thirdcategory and can be colored to indicate detected blood flow. Aspreviously mentioned, red can be used to represent blood flow towardsthe probe and blue can be used to represent flow away from the probe.Additionally, the intensity of the coloration can depict the strength ofthe backscattered Doppler signal. That is, when the ultrasound from theprobe goes through a “good” acoustic window at the skull, the blood flowsignal will appear intense and when the ultrasound goes through a “poor”acoustic window, the blood flow signal 152 will appear dull.

An example of an ultrasound system 600 (FIG. 6) according to anembodiment of the present invention that provides the features oflocating and tracking blood flow is described below. The particularembodiment will be described as utilizing a multi-element transducerhaving a six transducer elements. Those ordinarily skilled in the arthowever, will obtain sufficient understanding from the descriptionprovided herein to practice the invention with multi-element transducershaving greater or fewer transducer elements than described below to thespecific embodiment. The embodiment using the six-element transducerprovides an example which illustrates various aspects of the presentinvention that can be applied in alternative embodiments, such as thoseusing different multi-element transducers, using a different number oftransducer elements, or being utilized in different applications.

Those ordinarily skilled in the art will further appreciate that thenumerical processing employed to locate and track blood flow or tissuemotion, as described below, can be generalized to ultrasound systemshaving transducers with different numbers of elements. Thus, thesix-element transducer system described below for specific applicationssuch as observing blood flow in the MCA can be modified withoutdeparting from the scope of the present invention. Such modificationsinclude those made in light of geometric and physiologicalconsiderations. Examples of these considerations include the distancefrom the probe to the anatomical region to be observed, the size of thestructure to be observed, the lateral distance at depth of possiblelocations of the structure to be observed, the direction and speed ofthe motion to be observed, and the like. These considerations translateto design parameters such as array size, element size, pulse repetitionfrequency, number of look directions, ultrasound frequency, number ofpulses per look direction, and array apodization, as will be describedin more detail below.

FIG. 4A illustrates an ultrasound probe 410 having a multi-elementtransducer 400 according to an embodiment of the present invention. Thearrangement of transducer elements of the multi-element transducer 400is shown in FIG. 4B. A transducer as shown in FIG. 4B was constructed byBlatek (State College, Pa.). The transducer 400 includes six triangulartransducer elements 401-406 arranged in a hexagonal shape. The size ofthe hexagon is similar to the footprint of current single elementtranscranial Doppler probes, that is, measuring approximately 10 mmbetween two parallel sides of the hexagon. The hexagonal array isinscribed inside a circular package. All of the transducer elements arematched to 50 ohms real. Composite elements are used for the transducerelements to maximize sensitivity and minimize cross talk. A quarter wavematching layer is placed on the front of the array to maximizesensitivity on receive. A Faraday shield is placed around the entireprobe and connecting cable. The transducer 400 is capable of providingan unsteered transmit beam of 0.7 W/cm² (spatial peak temporal averageintensity using 16 cycle bursts at 8 kHz pulse repetition rate).

The transmit beam profile of the transducer 400 generally exhibitscircular symmetry and provides a 2.4 mm spotsize at the −3 dB intensitylevel at a depth of 50 mm. FIGS. 5A-5D illustrate delivery of a beamsteered to four different positions relative to a normal axis and havingbeam shape characteristics at the four different locations that arenearly identical. The beam is shown in FIGS. 5A-5 Dby the generallyconcentric circles, with each larger circle representing a region oflower beam intensity. The center of the beam can be electronicallypositioned up to 5 mm from the normal axis at 50 mm depth using a DSPplatform that is described in more detail below. The normal axis(extending perpendicular to the plane of the page with depth increasinginto the page) is identified by the intersection of the arrows in eachof the four figures. As shown in FIGS. 5A-5D, the ultrasound beam can besuccessfully steered to provide ultrasound in different directions.Using the beam steering capabilities, a plurality of sample locationscorresponding to look directions can be used to survey a region ofinterest. For example, steering the ultrasound beam can be used tointerrogate sample locations across the virtual surface 116 tointerrogate a region of interest.

FIG. 6 is a functional block diagram that depicts the Doppler ultrasoundsystem 600 in accordance with an embodiment of the invention. As will bedescribed in more detail below, the Doppler ultrasound system 600 can beused to electronically steer an ultrasound beam to a plurality oflocations in a region of interest to survey the region. For example, thevirtual surface 116 (FIGS. 1 and 2) can be constructed through beamsteering and Doppler shift signal processing. In this manner, blood flowin the region of interest can be located visually. In alternativeembodiments, a locating algorithm is executed to automatically locateblood flow in a region of interest, and in some embodiments,automatically tracked to maintain observation of the blood flow. Theultrasound system 600 includes a multi-channel Doppler DSP platform 700coupled to a probe 612 having a multi-element transducer (not shown inFIG. 6). The multi-element transducer includes a plurality of transducerelements with independent transmit and receive control of pulse packetphase and amplitude on each element. The probe 410 shown in FIG. 4A andhaving the arrangement of multi-element transducer 400 shown in FIG. 4Bcan be substituted for the probe 612. As will be described in moredetail below, the DSP platform 700 is configured to perform transmit andreceive beam steering by modifying phase and amplitude of the outgoingtransmit pulse of each active transducer element and analyzing thereceive signals from the steering direction. The receive function forthe DSP platform 700 will be described with respect to the presentembodiment as utilizing the same set of elements as used on transmit.However, alternative embodiments of the present invention are notlimited as such, and different numbers of elements can be used fortransmit and receive functions. The DSP platform 700 provides forgenerating transmit waveforms with variable amplitude and delay to drivethe transducer elements of the probe 612, digitizing receive echosignals detected by the transducer elements of the probe 612, and signalprocessing to generate Doppler shift data representing Doppler shiftsignals extracted from the receive echo signals.

The DSP platform 700 is coupled to a processing system 620 through a bus622. The bus 622 can be implemented using conventional computer bussesand protocols, for example, the bus 622 can be a universal serial bus(“USB”). The processing system 620 is configured for additionalprocessing of the Doppler shift data provided by the DSP platform 700and provides the DSP platform 700 with, among other things, commands anddata related to electronic beam steering of the transmit and receivesignals. Additionally, the processing system 620 executes algorithms forlocating and tracking blood flow, as described in more detail below. Theprocessing system 620 can be a host computer system to which the DSPplatform 700 is coupled, or alternatively, can represent processingsystems included in the DSP platform 700 or in an ultrasound system inwhich the DSP platform 700 is included for standalone Doppler signalprocessing, and locating and tracking of blood flow.

The processing system 620 is coupled to a display device 626 forproviding visual information and feedback to an operator. Theinformation can be displayed in different formats on the display device626. For example, for monitoring blood flow in the MCA, a format such asthat previously described with reference to FIGS. 1 and 2 can be used.The display device 626 can be a conventional display device now known orlater developed, including a flat panel display or cathode ray tube(“CRT”) display, which can be integrated with the ultrasound system 600,or is a standalone display device connected to the Doppler ultrasoundsystem 600 or to the processing system 620.

The processing system 620 is coupled to a data storage device 630 tostore data to or retrieve data from external storage media. Examples oftypical data storage devices 630 include hard and floppy disks, tapecassettes, compact disk read-only (“CD-ROMs”) and compact diskread-write (“CD-RW”) memories, and digital video disks (“DVDs”). Theultrasound system 600 is also coupled to audio speakers 632 forproviding audio information. The ultrasound 600 is further coupled to amicrophone 634 for receipt of audible information input by the user, andcoupled to one or more input devices 636, such as a keyboard or a mouse,to allow an operator to interface with the processing system 600.Although not shown in FIG. 6, the processing system can includeconventional circuits and software for storing the audio and visual datafor later playback and viewing.

FIG. 7 illustrates the DSP platform 700 according to an embodiment ofthe present invention. The DSP platform 700 includes sixTX/RX/DSP-channels 701-706, each of which is coupled to a respectivetransducer element of a six-element transducer 708 having transducerelements 708(1)-708(6). The transducer 708 is typically included as partof the probe 612 (FIG. 6). Each TX/RX/DSP-channel 701-706 includestransmit and receive circuits 710 and 712, respectively, coupled to aDSP 718 through logic circuit 716. The logic circuit 716 represents afront end processor that is used for performing repetitive tasks in theprocessing chain. It can be viewed as a link between the analog frontend of the TX/RX/DSP-channel represented by the transmit and receivecircuits 710 and 712, and the digital environment of the DSP 718. TheTX/RX/DSP-channels 701-706 are coupled through the data bus 622 to theprocessing system 620. As previously discussed, the processing system620 can represent a host computer system, processing systems includedwith the DSP platform 700, processing systems included in the ultrasoundsystem 600, or some other alternative processing system. An externalclock signal CLK and synchronization signal SYNC are provided to thelogic circuits 716 to synchronize operation of the logic circuits 716.The SYNC signal may be generated by a separate circuit, oralternatively, can be generated by the logic circuit 716 of one of thesix TX/RX/DSP-channels 701-706 and then passed to the other logiccircuits 716 for synchronizing operation of the remainingTX/RX/DSP-channels 701-706.

As previously discussed, the DSP platform 700 is coupled to amulti-element transducer 708, such as the transducer 400 (FIG. 4B). Eachtransducer element 708(1)-708(6) of transducer 708 has independenttransmit and receive control of pulse packet phase and amplitude. Inthis manner, a transmit beam delivered by the transducer 708 can beelectronically steered as shown in FIGS. 5A-5D. That is, the transmitbeam can be delivered along an ultrasound beam axis that is off-axisfrom a normal (i.e., perpendicular) axis of the face of the transducer708. As known, electronic beam steering can be accomplished by usingdifferent time delays and different amplitude weights (apodization) forthe waveform applied to each transducer element 708(1)-708(6) whendelivering a transmit pulse and receiving echoes. Timing and amplitudeof the waveform driving each transducer element 708(1)-708(6) iscontrolled by the respective TX/RX/DSP-channels 701-706 to performtransmit and receive steering. Note that beam steering is implemented onboth transmit and receive. On transmit, the transmit beam is directed toa target, and on receive, the receive signals are analyzed relative tothe steering direction.

FIGS. 8A and 8B illustrate the transmit circuit 710 and the receivecircuit 712 according to embodiments of the present invention. Withrespect to the transmit circuit 710 of FIG. 8A, the processing system620 provides commands for gain, phase (delay associated with steeringthe transmit beam), carrier frequency, length of the transmit burst andpulse repetition rate to the DSP 718 (FIG. 7). In one embodiment, eachtransducer is driven to deliver pulsed ultrasound having a carrierfrequency of 2 MHz and a pulse repetition frequency of 8 kHz. The DSP718 provides the appropriate digital commands to the logic circuit 716to activate a particular mode (i.e., transmit or receive) of operation.The logic circuit 716 generates two digital logic pulse trains V1 and V2with the specified carrier frequency (divider applied to externalclock), pulse repetition rate (initiated by external sync input), andpulse length. The two signals V1 and V2 are appropriately delayedrelative to the SYNC signal provided to the logic circuit 716 for therespective TX/RX/DSP-channels 701-706 in order to accomplish electronicbeam steering when all active elements are considered together. The twosignals are provided to an operational amplifier 820 included in thetransmit circuit 710, are added 180 degrees out of phase, and amplifiedby the operational amplifier 820 according to the gain specified by theprocessing system 620. The operational amplifier 820 is preferred toprovide approximately 40 dB of programmable transmit gain. However,amplifiers having other gain characteristics can be used as well. Thespecified gain is provided to the logic circuit 716 as digital data,which is converted by a digital-to-analog converter 822 into an analoggain signal applied to the operational amplifier 820. The resultingoutput signal from the operational amplifier 820 is a square “sinusoid”with center voltage of zero volts. The amplified signal is then appliedto a respective transducer element 708(1)-708(6) through a tuningcircuit (not shown in FIG. 8A) to drive the respective transducerelement 708(1)-708(6) to deliver transmit pulses. When combined, therespective transmit pulses of each transducer element 708(1)-708(6)results in a transmit beam delivered by the transducer 708.

With respect to the receive circuit 712 of FIG. 8B, receive signalsdetected by a respective transducer element 708(1)-708(6) are providedto a receive amplifier 824 preferably having a fixed gain ofapproximately 40 dB, and bandpass filtering that has center frequency at2 MHz and a bandwidth of approximately 300 kHz. In the present example,2 MHz is the center or carrier frequency for the transmit beam. Othercarrier frequencies and corresponding center frequencies can be used aswell. As understood by those ordinarily skilled in the art, modifyingthe carrier frequency may include consideration of the resulting beamgeometry, penetration through skull, and transducer elementconfiguration and size.

For each pulse period of ultrasound, echo signals resulting from thebandpass-filtering of receive signals are sampled at four times thecarrier frequency by an analog-to-digital converter (“ADC”) 826 toprovide digital echo data representing the echo signals. In the presentexample, the sampling frequency of 8 MHz. As shown in FIGS. 7 and 8,each channel includes the same functional blocks. In alternativeembodiments, some or all of the functional blocks can be combined intoone circuit that is shared or multiplexed for each of the differentchannels. For example, rather than having an ADC 826 for each channel, asingle ADC having sufficient sampling rate can be shared by all of thechannels for digitizing echo signals received by the respectivechannels. Other functional blocks can be combined as well withoutdeparting from the scope of the invention.

The echo data are processed by the logic circuit 716 to demodulate theecho signals from a pulse period of ultrasound into Doppler (I,Q) shiftsamples that stratify the depth range of interest along the ultrasoundbeam. As known in the art the “I” value represents a measure of aDoppler shift sample along an “in-phase” or “real” axis of the complexplane and the “Q” value represents a measure of the Doppler shift sampleat essentially the same time and position, but on a “quadrature” or“imaginary” axis of the complex plane. In addition to the Doppler shiftsamples, the echo data generated by the ADC 826 are also output by thelogic circuit 716. As will be explained in more detail below, thedigital echo data, along with the Doppler shift samples, can beprocessed for identifying a direction for monitoring blood flow.

The Doppler (I,Q) shift samples and the echo data generated by the logiccircuit 716 are provided to respective DSPs 718 (FIG. 7), whichconstructs Doppler shift signals from multiple Doppler shift samples.Each Doppler shift signal is constructed from Doppler shift samples fromthe same echo depth and across multiple pulse periods. The Doppler shiftsignals are represented by Doppler shift data that are output by therespective DSPs 718.

FIG. 9 is a flow diagram illustrating the signal and data processingperformed by the DSP platform 700 for receive signals. The receivesignals are amplified and band-pass filtered at the carrier frequency atstep 902. The band-pass filtered signal is then digitized to providedigital echo data representing the receive signal at step 904. The echodata is demodulated at step 906 to generate Doppler (I,Q) shift samplesfor the depth range that are low-pass filtered at step 908 to removenoise outside the bandwidth of the transmit signal. Demodulation andlow-pass filtering are performed within each pulse period. At step 910,the low-pass filtered Doppler (I,Q) shift samples are decimated to carryforward only candidate signals at the depth or at specific depths acrossa depth range of interest. Doppler shift signals represented by Dopplershift data are constructed at step 912 from the Doppler shift samplesfrom multiple pulse periods. The Doppler shift signals are clutterfiltered at step 914 to suppress “clutter” from each Doppler shiftsignal which can introduce severe artifact into a locating algorithmperformed for identifying blood flow.

Suitable methods for demodulating, low-pass filtering and decimating aredescribed in greater detail in the aforementioned U.S. Pat. No.6,196,972 to Moehring. However, it will be appreciated that othermethods can be used as well. In summary, the aforementioned patentdescribes a demodulation process for generating Doppler (I,Q) shiftsamples that can be performed through simple subtraction operationsoperating on each successive quartet of samples of echo data for a pulseperiod. Each sample corresponds to digital echo data resulting fromsampling the signal output by the receive amplifier 824 at four-timesthe carrier frequency. For a quartet of samples, the third value issubtracted from the first value to produce the real part “I,” and thesecond value is subtracted from the fourth value to produce theimaginary part “Q” of a complex Doppler (I,Q) shift sample for anassociated depth. The same operation is performed on all quartets ofsamples of echo data for a pulse period, with each succeeding quartet ofpoints associated with a location of greater depth. The particulardemodulation method blurs the axial resolution by approximately onewavelength of the carrier, but is acceptable in typical applicationssince one wavelength of the carrier is inconsequential relative to thetypical sample volume size associated with medical pulse Dopplerultrasound.

Demodulation into Doppler (I,Q) shift samples is followed by a low-passfilter operation. The low-pass filter operation described in theaforementioned patent involves taking as many as 35 contiguous gatepositions bracketing a desired gate depth, within one pulse period, andapplying a low-pass finite impulse response (“FIR”) filter. The processof low-pass filtering reduces out-of-band noise from a signal which issampled across successive pulses at a relatively low frequency (i.e.,the pulse repetition rate). The FIR filter is applied to the Doppler(I,Q) shift samples spanning the depth range bracketing a gate toconstruct one Doppler (I,Q) shift sample for each particular gate forthe particular pulse period. The desired gate depth or range of gatedepths that are output from this process will depend on the applicationof the Doppler ultrasound system. For example, for monitoring blood flowin the cerebral arteries, the desired gate depth is approximately 50 mmfrom the probe, with a bracketing gate depth range of 40 mm to 60 mm.However, other applications of the Doppler ultrasound system may dictatea different range of Doppler gate computation.

Clutter cancellation in the signal location mode outlined here isaccomplished with FIR filters due to the short sequences of Dopplershift signals anticipated in order to maximize signal acquisition speed.Infinite impulse response (“IIR”) filters are appropriate in thetracking mode where adjustments to the transmit and receive steering areless frequent and the sequences are significantly longer. A simplederivative filter can also be used to clutter filter the low-passfiltered data, across multiple pulse periods and at a fixed depth. Forexample, in a set of time series data, each pair of adjacent samples canbe subtracted to produce a clutter-filtered value in the time seriesdata. The clutter filtered Doppler shift data is then provided to theprocessing system 620 from each of the six TX/RX/DSP-channels 701-706.The data can be used by the processing system 620 to construct an imagefor display on the display 626 (FIG. 6) to provide visual feedback to anoperator for the presence of blood flow in a region of interestinterrogated by ultrasound at a plurality of sample locations. Forexample, using the Doppler shift data provided to it, the processingsystem 620 can construct an image of a compound mode Doppler virtualsurface 116 (FIG. 2), as previously described. Although not discussedherein in detail, construction of an image from the Doppler shift datacan be accomplished using conventional techniques now known by thoseordinarily skilled in the art, or later developed. Consequently, in theinterest of brevity, a detailed discussion of constructing such an imageis omitted from herein. As will be explained in more detail below, theDoppler shift data can be further processed by the processing system 620according to a locating algorithm to locate blood flow in a region ofinterest.

With the use of a locating algorithm, the DSP platform 700 and theprocessing system 620 can be utilized to quickly and automaticallyidentify (i.e., locate) blood flow in a region of interest, and continueto monitor (i.e., track) the blood flow without the need for manuallyadjusting the location or orientation of the ultrasound probe. As willbe explained in more detail below, the locating algorithm applies modernspectral estimation theory previously applied in seismic array frequencyand wave number analysis for locating blood flow. More specifically, theminimum variance spectral estimation technique of Capon is applied tothe collection of Doppler shift signals, each Doppler shift signalderived from a respective transducer element, for automatically locatingand tracking blood flow or tissue motion.

The application of the minimum variance spectral estimation technique ofCapon for the locating algorithm is described with respect to the sixelement transducer 400. This description will pertain to increasingdetail in vicinity of one transmit direction (i.e., look direction)among all those used to tile the virtual surface 116 (FIG. 2). Thisdiscussion applies to the spatial region corresponding to the lateralextent of the transmit beam about the given look direction, and at adepth of interest. Reflections outside this region will not contributesignificantly, and therefore, the corresponding “out of beam” signalprocessing is not performed. As previously discussed, the use of the sixelement transducer 400 provides a relatively simple vehicle to describeapplication of the algorithm. However, those ordinarily skilled in theart will obtain sufficient understanding from the description providedherein to modify the locating algorithm and tracking algorithm forapplication with other multi-element transducers having greater or fewertransducer elements, or having the transducer elements arrangeddifferently.

Generally, the embodiment of the locating algorithm and trackingalgorithm described herein is capable of determining locations of onefewer signal sources than the number of individual elements in thetransducer. The signals here are those intersecting a spherical surfacewhich is a constant distance, D, away from the transducer 400, such asthe virtual surface 116 (FIG. 1A). In the particular application of TCD,where the signals of interest represent blood flow in the basal arteriesin the brain, including the middle cerebral artery, the anteriorcerebral artery, and the posterior cerebral artery, the signals are farapart relative to the transmit beam width. Therefore, there is generallyone signal in the evaluation neighborhood about the given lookdirection. Regarding any of the basal arteries of the brain, the numberof vessels in the search region and at a typical depth is one. As aresult, there is very straight forward positive confirmation to performonce a signal is discovered, because generally finding more than onesignal in the region of interest will be unusual.

The geometric center of each transducer element 401-406 (FIG. 4) will bereferred to as the “phase center”, {right arrow over (c)}_(i)=(c_(xi),c_(yi)), where i is the index of the transducer element (1 through 6).As previously discussed, the receive signals are digitized at four- oreight-times the carrier frequency, the resulting data demodulated,low-pass filtered, decimated, and clutter filtered so that basebandDoppler shift samples can be constructed for the depth D for eachelement and for each pulse period. After clutter filtering the samplesfor a fixed-gate depth, a variable α_(ij) is defined as the complexDoppler shift sample associated with the i^(th) transducer element andthe j^(th) pulse period of the Doppler signal (at depth=D). For a fixedi, α_(ij) represents a Doppler shift signal, whereas in total, α_(ij)represents “Doppler shift data.” The Doppler shift data collected fromdepth D over N pulse periods are expressed as

p _(j=1 . . . N)=[α_(1j),α_(2j),α_(3j),α_(4j),α_(5j),α_(6j)]^(T)  (0.1)

where T denotes non-conjugated transpose. A covariance matrix of thisprocess contains the information of interest regarding the location ofblood flow or tissue motion. The covariance matrix is calculated by

$\begin{matrix}{{R = {{{cov}(p)} = {\frac{1}{N}{\sum\limits_{j}\; {p_{j}p_{j}^{H}}}}}},} & (0.2)\end{matrix}$

where H indicates conjugate transpose.

The concept of a “steering vector” will now be introduced. The steeringvector is used to search across the solid angle at depth D for bloodflow or tissue motion. For the purpose of providing a common frame ofreference, it is assumed that the transducer is positioned in an x-yplane at a depth z=0. A plane wave moving across the transducer aimedwith arbitrary spherical coordinate angles θ and φ will have phase attransducer element i expressed as

$\begin{matrix}{\Phi_{i} = {^{j\; {\overset{\rightarrow}{k} \cdot {\overset{\rightarrow}{c}}_{i}}} = {\exp \left\lbrack {j\frac{2\pi}{\lambda}\left( {{c_{xi}\cos \; {\theta sin\phi}} + {c_{yi}\sin \; {\theta sin\phi}}} \right)} \right\rbrack}}} & (0.3)\end{matrix}$

where {right arrow over (c)}_(i)=(c_(xi), c_(yi)) is the phase centerfor the element and

$\overset{\rightarrow}{k} = {\frac{2\pi}{\lambda}\left( {{\cos \; \theta},{\sin \; \theta}} \right)\sin \; {\phi.}}$

The symbol φ in these expressions represents the angle from the z-axisand runs from 0 to π, and the symbol θ represents the angle in the x-yplane, measured from the x-axis, and runs from −π to π. A plane wavetraveling in the ({circumflex over (θ)}, {circumflex over (φ)})direction will have a different phase, Φ_(i) (θ,φ), associated with eachtransducer element. Note that the beam direction for a steering vectoris established by choosing θ and φ and calculating Φ_(i) (θ,φ)associated with each transducer element. The associated phases for allelements comprises the “steering vector”:

s(θ,φ)=[(Φ₁(θ,φ),Φ₂(θ,φ),Φ₃(θ,φ),Φ₄(θ,φ),Φ₅(θ,φ),Φ₆(θ,φ)]^(T)  (0.4)

The steering vector specified by the independent variables θ and φ isused to explore the signal power received from the transducer fromdifferent directions.

A minimum variance power spectrum, which indicates Doppler signal poweras a function of steering direction, is expressed as a quadraticfunction of the steering vector and the covariance matrix inverse:

$\begin{matrix}{{P^{MV}\left( {\theta,\phi} \right)} = \frac{1}{{S^{H}\left( {\theta,\phi} \right)}{R_{s}^{- 1}\left( {\theta,\phi} \right)}}} & (0.5)\end{matrix}$

Once the region around a normal axis is explored by varying θ and φ, aparticular location will be identified as corresponding to the locationof blood flow or tissue motion and a steering vector defined for thelocation, or the information from varying θ and φ will be used to fillin local flow details in the virtual surface 116 of FIG. 1B.

Apodization weights and delays to apply to the six transducer elementsto steer transmit and receive signals can then be specified by the sixelement vector w:

w=R ⁻¹ s({circumflex over (θ)},{circumflex over (φ)})P ^(MV)({circumflexover (θ)},{circumflex over (φ)}).  (0.6)

Here ({circumflex over (θ)},{circumflex over (φ)}) is the location ofthe maximum calculated Doppler signal power, |P^(MV)|, over the regionprobed with the steering vector and further probed with varying transmitdirection. To electronically steer the ultrasound beam on eithertransmit or receive, the array apodization weight values are themagnitudes of the elements of w which are applied to each transducerelement as gain coefficients, and the individual element time delays areobtained from the phases of the elements of w divided by 2πf₀.

Operation of the Doppler ultrasound system 600 and the processing system620 will now be described with specific application to TCD. Aspreviously discussed, in locating cerebral blood flow, an acousticwindow in the skull through which the blood flow can be observed by theDoppler ultrasound system is initially located. The desired blood flowis then located and monitored. A method for using the DSP platform 700and transducer 400 to locate an acoustic window will be described below,followed by a description of a locating algorithm according to anembodiment of the invention. In addition to the DSP platform, theprocessing system 620 (FIG. 1) is used in executing the locatingalgorithm and performing the process of locating an acoustic window. Itwill be appreciated by those ordinarily skilled in the art, however,alternative embodiments of the invention may include additionalprocessing sub-systems or combine some or all of the processingcapabilities of the processing system 620 in the DSP platform. Thus, thescope of the invention is not limited to a particular allocation ofprocessing between the DSP platform and the processing system 620.

FIG. 10 is a flowchart illustrating a process of identifying an acousticwindow according to an embodiment of the present invention. An acousticwindow is identified as a probe location from which blood flow or tissuemotion can be observed using ultrasound. For example, with respect to aspecific application of monitoring cerebral blood flow, the acousticwindow is the location on the skull that allows one to acoustically“see” through the skull bone to observe blood flow. In finding anacoustic window, the ultrasound probe is positioned on the skull. Atstep 1002, transmit signals are delivered from the current position.Receive signals, resulting from reflected transmit signals, aremonitored for a depth range of interest. In the specific example oflocating cerebral blood flow of the middle cerebral artery, the depthrange of interest is generally between 40 and 60 mm with a particulardepth of interest at 50 mm. At step 1004 the amplitude of a group ofreceive signals are averaged and at a step 1006 are compared to a firstthreshold value to determine whether there is a lack of receive signalsfrom the depth range of interest. In one embodiment, raw digital echodata can be used in generating data that is compared to the firstthreshold value. For example, a sum of squared amplitudes can begenerated from the digitized samples with the summation starting andending at specific depths (e.g., 40-60 mm). In another embodiment, theamplitude of each successive group of 16 ultrasound pulses are averaged.For example, an average absolute amplitude can be computed, again over adepth range, and then compared with an amplitude threshold.

If the average absolute amplitude is less than the first thresholdvalue, it is assumed that the probe should be relocated since thereflections from the present probe location are not sufficient toindicate backscatter from underlying brain tissue, and therefore, theprobe is not positioned over an acoustic window. At step 1008, the useris provided with feedback that the probe should be moved to a newlocation. If the average amplitude at the step 1006 is greater than thefirst threshold value, then at step 1010 a second threshold used toevaluate whether there is a Doppler signal somewhere in the field ofview of the probe. The Doppler signal power can be calculated by summingthe squared absolute values of the Doppler shift signal, and thenfurther summed over a depth range of interest and compared to the secondthreshold. If the Doppler signal power is less than the secondthreshold, feedback is provided at stop 1012 to the user that the probeis near or over an acoustic window but there is no detected blood flowsignal in the viewing range of the probe. Accordingly, the position ofthe probe should be changed with respect to the current location, whichincludes changing the direction of the probe but not the position on theskull. However, if a Doppler signal is detected, that is, the Dopplersignal power is greater than the second threshold, then a second form offeedback is given to the user at step 1014 indicating that an acousticwindow and underlying blood flow has been located.

In providing feedback on whether the average amplitude of the receivesignals suggests the presence or absence of an acoustic window for aparticular probe location, and further whether there is underlying bloodflow, audio and/or visual feedback can be provided to assist theoperator in finding an acoustic window. For example, audio feedback inthe form of tone having a variable volume that changes as the probe ismoved towards or away from an acoustic window can be used. Visualfeedback in the form of an image on a graphical display can supplementor replace audio feedback. In one embodiment of the invention, a processanalogous to a “stud finder” can be used. That is, the probe can bemoved over the surface until a light emitting diode (“LED”) emits light,indicating that the probe is over an acoustic window. The operator canthen adjust probe motions to be angular (“flashlighting”) rather thanlateral motions across the temporal bone region, in searching for bloodflow. When blood flow signals are located, a second LED may be employedto convey this information to the operator, indicating that thetransducer should be secured in a fixed position over the currentlocation. Other forms of feedback known in the art can be used as well.

After an acoustic window has been located, and the transducer ispositioned accordingly, and a locating algorithm can be performed forautomatically locating and tracking blood flow or tissue motion. Analgorithm for automatically locating and tracking blood flow accordingto an embodiment of the present invention will now be described.

The algorithm has two “modes” of operation while performing theauto-location of blood flow and tracking of the same. The two modes canbe generally described as a search mode, and an acquisition/track mode.In the search mode, blood flow in a region of interest is located bybeam steering transmit and receive signals over a search region. Anexample of a search region is provided by a compound-mode Dopplervirtual surface 116 having a plurality of sample locations that areinterrogated with ultrasound. Based on the information obtained for theplurality of sample locations, blood flow is located. Inacquisition/track mode, the located blood flow is acquired andprocessed. A set of apodization weights and delays, represented by asteering vector, are applied to the individual elements of amulti-element transducer, such as the transducer 400, in locating,acquiring and tracking the acoustic reflector of interest. The set ofweights and delays determines the form and direction of the ultrasoundthat is delivered by the transducer. Steering vectors which accomplishelectronic beam steering across two degrees of angular freedom are bothutilized and derived in search mode to determine where blood flow islocated. Steering vectors may also be utilized in different fashion. Ageneral “raster scan” search may be employed to search in a brute forcefashion for blood flow in a region of interest. A second type of searchmay employ signal processing search algorithms such as a gradient searchor a simplex search, and more quickly accomplish target location. Theselatter techniques may be employed with broader beam shapes, that is,beams that are significantly larger than the target and havingsubstantial size regarding the search region. These sorts of beams willbe more amenable to giving steering feedback when the maximum intensityportion of the beam is pointed away from the target, but the targetstill rests inside the mainlobe of the beam. A broader beam may beformed by apodization (both amplitude and phase) of the transducerelements, as known in the art.

The steering vector identifying blood flow in the region of interest isthen carried into the acquisition/track mode of operation. The locatingalgorithm generally uses a fixed steering vector in acquisition/trackmode. Updates are made to the steering vector in acquisition/track modebased on incidental changes in the location of the blood flow relativeto the probe, which may be instigated by a variety of phenomena. Forexample, the changes can be caused by the patient moving his head andthe probe being blocked from following this motion, the probe beingjarred into a new position by a surgeon's elbow, and pronouncedmastication. Changes in the location of blood flow can be sensed bycontinually evaluating the local neighborhood of the transmit beam lookdirection associated with the acoustic reflector of interest, orintermittently scanning a broader region about the transmit beam lookdirection associated with the location of blood flow. In either method,an updated look direction is established for the acoustic reflector ofinterest.

In alternative embodiments of the invention, the spacing and number ofsample locations interrogated in the region of interest can be modifiedaccording to the mode of operation of the locating algorithm. That is,during search mode, to survey a large region of interest, the number ofsample locations and the spacing of the sample locations can be selectedto interrogate a large region of interest. When blood flow is acquiredfrom executing the locating algorithm, and the ultrasound is steered inthe direction of the location of the blood flow, the number of samplelocations and/or the spacing of the sample locations in the regionsurrounding the location identified for the blood flow can be changedfrom the number and spacing of the sample locations used during searchmode. In this manner, modifying the number and spacing of the samplelocations from one mode to another can be advantageously used toefficiently search a region in search mode, and then used to obtain highresolution information from a region adjacent the located acousticreflector of interest. More generally, changing the number of samplelocations, as well as the spacing or density of sample locations is amodification that can be made without departing from the scope of thepresent invention.

As previously discussed, blood flow at a depth of interest in a regionof interest is initially located during the search mode. In the searchmode, the ultrasound transmit and receive signals are steered to aplurality of look directions over a search region, and information isacquired for a depth range bracketing the depth of interest for eachcorresponding sample location. In one embodiment, the sample locationscorrespond to a respective look direction across a compound-mode Dopplervirtual surface 116, as previously discussed with respect to FIG. 2.Information at several different depths in the depth range can be takenfor each look direction. The depth range bracketing the depth ofinterest can span approximately 20 mm and include between five and 10different depths. Data is acquired from each depth. As will be explainedin more detail below, the information acquired at the different depthswill be used to confirm that a potential blood flow signal is not anartifact or spurious noise. The locations and the depth range for whichinformation is acquired can be preset to provide a scanning pattern thatfacilitates locating blood flow. Various factors can be considered andbalanced in establishing a scanning pattern. For example, the size ofthe search region, the number of sample locations at which informationis acquired, the desired resolution to locate blood flow in the searchregion. Additionally the number of sample locations should be balancedagainst the processing overhead of the DSP platform 700 and theprocessing system 620 in which the locating algorithm is executed.

For each of the look directions and for the different depths in thedepth range bracketing the depth of interest at which information isacquired, the Doppler signal strength of the receive signal from therespective location and depth is calculated. Conventional algorithms canbe used for the calculation of the Doppler signal strength for thedifferent locations and the different depths. Previous discussionsdescribing the calculation of Doppler signal power can also be applied.An example of a suitable process for calculating Doppler signal strengthis described in the aforementioned U.S. Pat. No. 6,196,972 to Moehring,which has been incorporated herein by reference. The power valuescalculated for the different locations and depths form an array of datafrom which a location corresponding to a maximum calculated power for adepth of interest can be determined. The identified location representsthe location of blood flow in the region of interest. A steering vectorcorresponding to the direction of the identified location is calculatedand a set of apodization weights and delays are resolved forelectronically beam steering transmit and receive signals in thedirection of the blood flow.

The location of blood flow can be confirmed by comparing the calculatedDoppler signal strength for the location at the depth of interest withthe calculated Doppler signal strength in the same look direction for atleast one other depth in the depth range. Confirmation is made if thecalculated Doppler signal strength for the same look direction but atthe different depth or depths is a value that is consistent with a bloodflow signal that is present at multiple depths. In contrast, if thecalculated Doppler signal strength for the same look direction anddifferent depth indicates that a sufficient Doppler signal strength ispresent only for the depth of interest, it is unlikely that the locationcorresponds to blood flow in the region of interest.

After a location for the blood flow is identified in the search mode,the ultrasound system enters an acquisition/track mode during which thesteering vector identifying the location of the blood flow at the depthof interest is updated by applying the minimum variance spectralestimation technique previously described. A new location correspondingto the peak of the minimum variance power spectrum provides the updateddirection for electronically beam steering the transmit and receivesignals. Specifically, the updated steering vector is determined bycalculating a power spectrum for a region proximate to the “old”location of the acoustic reflector of interest by varying the directionof the old steering vector. As previously discussed, the direction ofthe steering vector can be defined by two spherical coordinate angles, θand φ, relative to the transducer. The θ and φ corresponding to themaximum calculated power over the region proximate the old location isused to identify the updated steering vector defining an updatedlocation of blood flow. Based on the updated steering vector that iscalculated, the appropriate apodization weights and delays are resolvedand applied to steer the transmit and receive signals in the directionof the updated location. With the updated location of blood flowidentified and the apodization weights and delays determined, thetransmit and receive signals are electronically steered to monitor theblood flow by acquiring and processing data from the updated location.

The minimum variance power spectrum can be calculated concurrently withthe monitoring activity, which minimizes any discontinuities in theinformation that may occur in updating the transmit and receivebeamformers with the new direction. It will be appreciated that theupdate will result in a discontinuity in the Doppler shift signal, andtherefore, cannot be updated continually due to spectral analysisartifacts. As a result, updating can be performed every one or twoseconds instead, or alternatively, whenever a new location is greaterthan a threshold distance from the old location.

In the event the position of the probe has shifted to such a degree thatthe blood flow signals are no longer in the region proximate to thecurrent steering vector, the search mode is reentered to again locatethe blood flow. In one embodiment, the maximum calculated power of thepower spectrum is compared to a minimum threshold value to determine ifthis occurs. The minimum threshold value should be selected such that acalculated maximum power value below the threshold value is a goodindication that the probe position has shifted enough so that thetransmit and receive signals are no longer aimed in the generaldirection of the location of blood flow. If the maximum calculated poweris below the minimum threshold, then it can be assumed that the bloodflow is no longer within the search region probed by adjusting the θ andφ of the current steering vector. In response, the search mode isreentered and the process of locating blood flow, followed byacquisition and processing of data from the direction of the locatedblood flow, is performed as previously described.

In summary, after positioning a probe having the multi-elementtransducer, such as the transducer 400, over an acoustic window, thepreviously described algorithm can be utilized to quickly andautomatically locate blood flow in a search region and continue tomonitor the blood flow. FIG. 11 is a flow diagram illustrating thegeneral process of the locating algorithm according to an embodiment ofthe present invention. At step 1102, for each sample location sampled,Doppler shift signals are extracted from the echo signals. An example ofextracting Doppler shift signals was previously described. At step 1104,Doppler signal strength for the sample locations is calculated. Based onthe Doppler signal strength calculation at step 1104, it is determinedwhether an acoustic reflector of interest is located in the searchregion at step 1106. At step 1108, if the power calculation does notindicate blood flow in the current search region, the ultrasound probeis repositioned and a new search region is interrogated at step 1109.The process returns to step 1102 to begin extracting Doppler shiftsignals from the echo signals of the new search region. If, however,blood flow is identified at step 1108, a steering vector is calculatedfor the location of the blood flow at step 1110. Using the steeringvector, apodization weights and delays are calculated to electronicallysteer transmit and receive signals in the direction of the location ofblood flow at step 1112. Following step 1112, the blood flow has beenacquired and can be monitored.

Tracking of the blood flow generally begins at step 1114, where a regionproximate to the location of the blood flow is interrogated withultrasound and an angular power spectrum is calculated for theinterrogated region. At step 1116 a minimum variance spectral estimationtechnique is applied to the angular power spectrum to identify alocation of peak power corresponding to an updated location for theblood flow. At step 1118, if the maximum calculated power from step 1116is no longer indicative of blood flow located in the proximate regioninterrogated at step 1114, suggesting that the probe has significantlyshifted and the blood flow will need to be reacquired, a search mode isreentered through step 1109 to perform a search of a search region thatis broader than the region interrogated at step 1114.

However, if at step 1118 the maximum calculated power does indicate thatblood flow is in the region interrogated, and the updated location ofthe blood flow is sufficiently distant from the former location of theblood flow, an updated steering vector for the updated location of theblood flow (corresponding to the location of peak calculated power) iscalculated at step 1120. At step 1122, apodization weights and delaysare calculated to electronically steer transmit and receive signals inthe direction of the updated location of the blood flow. At step 1124,if the monitoring of the blood flow is to continue, the tracking processbegins again by returning to step 1114. Otherwise, the tracking processterminates.

As previously discussed, in the specific application of observing bloodflow in the vicinity of the proximal MCA, although embodiments of thepresent invention directed to the specific application are not limitedto such, it is preferable for a virtual surface 116 (FIG. 2) to have adiameter of 20 mm or greater. As also previously mentioned, beamsteerability and beam width are factors that respectively affect thediameter of the virtual surface 116 and the number of look directionsused to tile the area with sufficient density for locating blood flow inthe search region. These factors are in turn determined by element size,gross array dimensions, and apodization. For example, as known in theart, generally, the larger the element size, the lower the steerability.Additionally, the larger the gross array aperture dimensions, thenarrower the resulting beam. As a result, delivering ultrasound having arelatively narrow beam may dictate including space between adjacentelements (given a limited number of elements to utilize), which in turnraises the issue of “grating lobes.” Moreover, in deciding thearrangement and number of transducer elements to include in amulti-element transducer array, some practical considerations should bemade. For example, the number of transducer elements may be limited bythe complexity and the cost of the associated electronics associatedwith operating such an ultrasound transducer. Thus, as demonstrated bythe specific example, in designing a multi-transducer array for aspecific application, several factors, such as the number andarrangement of transducer element, should be considered.

FIG. 12 illustrates various relationships between beam steeringcapability, number of elements, and transducer element size, as relatedto total active area of a 13 mm diameter probe surface delivering 2 MHzpulsed ultrasound. The beam steering capability is represented by the“viewing diameter” of a virtual surface 116 for a depth of 50 mm fromthe probe surface. As shown in FIG. 5, a suitable element size for theapplication of monitoring blood flow of the MCA, that is, at a depth of50 mm and a virtual surface of 20 to 25 mm in diameter, is between 2.0and 2.5 mm to provide sufficient beam steering. As previously mentionedin order to maintain “entrance beam dimensions” at the skullcorresponding to a safe and reasonable heating level of the temporalbone (thermal index cranial, TIC<2), an ultrasound probe having adiameter of 13 mm is selected. The probe size results in a specificinter-element spacing for a given number of elements, which also resultsin “dead space” between elements. The dead space effectively moderatesthe active area of the probe. The right-hand axis of FIG. 12 shows thepercentage active area and the three curves correspond to differentnumbers of active transducer elements used for generating the ultrasoundbeam. More specifically, the three curves correspond to deliveringultrasound from 7, 19 and 37 active elements.

Based on the criteria previously described for the application ofmonitoring blood flow of the MCA, FIG. 12 suggests that a 19 elementprobe is preferable. The resulting 19 element probe has about 60% activearea compared to a solid piston having a 13 mm diameter. Although the 19elements compose only 60% of the 13 mm diameter, the voltage applied tothis probe can be increased to moderate the effect of deliveringultrasound from a reduced active area.

FIGS. 13A-13C illustrate multi-element transducers 1300, 1350, and 1370according to alternative embodiments of the present invention. Themulti-element transducers 1300, 1350, and 1370 are examples ofmulti-element transducers that are designed considering factorspreviously discussed with reference to FIG. 12. The individual elementsof the transducers are foreseen to have lower side lobe activity andgreater circular symmetry when compared to the triangular elements ofFIG. 5B. Additionally, the individual elements are smaller, and as aresult, will broadcast a spatially wider beam (+/−10 mm laterally at 50mm depth) in comparison to the elements of the transducer 400. Thisenables a greater degree of beam steering. However, a trade off withusing a greater number of smaller elements is the cost of controllingthem, which can be costly both in design efforts and in resultinginstrumentation size, and decreased power output from any given element.

The transducers 1300, 1350, and 1370 have 19 active transducer elements.In the transducer 1300, the 19 elements are configured as shown in FIG.10A, and are used without any additional transducer elements. Incontrast, the transducers 1350 and 1370 have an array of transducerelements, and 19 transducer elements of the transducers 1350 and 1370are used for an “active” transducer area. The 19 element footprint maybe shifted around the total footprint of the transducers 1350 and 1370during the process of locating and tracking signals of interest.

The transducers 1300, 1350, and 1370 use the same basic elemental tile,which is a hexagon with 2 to 2.5 mm distance between sides (notvertices). The ultrasound transducers 1300, 1350, and 1370 preferablyhave each transducer element tuned to 50 ohms, a cable that will notexceed about 5 mm diameter, quarter wave matching layer, 2 MHz carrierfrequency, minimum 20 percent bandwidth, cross talk not to exceed −30 dBbetween adjacent elements, electrical cross talk to be 5 to 10 dB lowerthan mechanical cross talk, a 200 micron kerf, and a Faraday shieldaround the entire probe which includes the cabling back to chassis.

The transducers 1300, 1350, 1370 designed with the idea of minimizingoperator requirements for moving the probe around the temporal boneregion while the system samples for underlying blood flow in TCDapplications. The transducer 1370 can be placed with its longitudinalaxis parallel to a line drawn between the ear and the eye, and thenmoved around the temporal bone region. A local transducer group of 19active transducer elements are shown in bold lines in FIGS. 10B and 10C.It will be appreciated that if the local transducer group is translatedthe width of one element to the right, certain new elements will becomepart of the active transducer area and others will no longer be part ofthe active transducer area. There are some advantages to having thelocal transducer footprint as small as the outlined 19 element regionsin FIGS. 10B and 10C, or even smaller. For example, the smallertransducer footprint provides a broader transmitted beam, and willresult in fewer “interrogations”—outgoing pulses—for a search region.However, a necessary consideration of the size of the local transducerfootprint is the possible increase in thermal index cranial.

The transducers 1300, 1350, and 1370 can be utilized with a DSP platformthat is capable of driving at least 19 transducer elementsindependently, and to receive and process receive signals detected by atleast 19 transducers. Although the DSP platform 700 of FIG. 7 wasdescribed as including only 6 independently controlled TX/RX/DSPchannels 701-706, those ordinarily skilled in the art will obtainsufficient understanding from the description provided herein to providea DSP platform that can be used with the transducers 1300, 1350, and1370. Such modifications remain within the scope of the presentinvention.

Although FIG. 12 suggests using 19 active transducer elements ispreferable for the specific application of monitoring blood flow in theMCA, and transducers 1300, 1350, and 1370 are shown having 19 activetransducer elements, the number of active transducers, the elementaltile shown, and the arrangement and number of transducer elements in anarray, have been provided by way of example. The previous description isdirected to particular embodiments of the invention, and are notintended to describe limitations limiting the scope of the presentinvention.

Turning to an alternative embodiment of the invention, the functionalitypreviously described are combined to provide automatic detection andtracking of blood flow with assessment of characteristics of the bloodflow. FIGS. 14A-14D illustrate various images that are displayed to anoperator during operation of an embodiment of the present invention thatprovides the combined functionality. Some or all of the principlespreviously described can be applied in the present embodiment of theinvention. As previously discussed, those ordinarily skilled in the artwill obtain sufficient understanding of the invention from thedescription provided herein to practice the invention, including thepresent alternative embodiment.

FIG. 14A illustrates an image of a virtual surface 1416 for a firstposition of an ultrasound probe on a patient's skull. As previouslydescribed with respect to the virtual surface 116 (FIG. 2), a pluralityof sample locations are interrogated, and information obtained from eachof the sample locations can be used to construct the virtual surface1416. The image has a shadow region 1420 on the right hand side of theimage of the virtual surface 1416. The portion of the virtual surface1416 corresponding to the shadow region 1420 represents the regionultrasound cannot penetrate bone of the skull, that is, a region notaligned with an acoustic window. In contrast, the left side of the imageof the virtual surface 116 is partially covered with gray signals forthe look directions 210. The gray coloration indicates that theultrasound can penetrate the bone for the corresponding look directions210, but that the reflections being observed for the depth of interestdo not contain Doppler information indicating blood flow. Theinformation for constructing the image of the virtual surface 1416 canbe obtained through the acoustic window locating algorithm previouslydescribed, as well as the construction of Doppler shift information forthe plurality of sample locations in the region of interest, representedin FIG. 14A as a compound mode Doppler virtual surface 1416.

FIG. 14B illustrates an image of the virtual surface 1416 for a secondposition of the ultrasound probe on the patient's skull. An acousticwindow locating algorithm providing sufficient information to constructthe visual feedback image enables an operator to move the probe from thefirst location illustrated in FIG. 14A to an acceptable location over anacoustic window, as shown in FIG. 14B. In contrast to FIG. 14A, theimage of the virtual surface 1416 of FIG. 4B does not have any shadowregions 1420, indicating that all of the look directions 210 across thevirtual surface 1416 are aimed through an acoustic window. A locatingalgorithm is executed to process the data obtained for the plurality ofsample locations in order to locate blood flow, as previously described.Additionally, using the Doppler shift signals extracted from the echosignals from each of the sample locations, blood flow information can begenerated. A red signal (not shown in color in FIG. 14B) is displayedfor a look direction 1430. The red signal represents a Doppler bloodflow signal 1432 automatically detected for the look direction 1430 andat the depth of interest. The blood flow signal 1432 will remaindisplayed as long as blood flow is detected for the look direction 1430and the probe remains in the same position. The blood flow signal 1432can be further visually highlighted to alert the operator that a flowsignal has been detected. As shown in FIG. 14B, the blood flow signal1432 in the look direction 1430 is highlighted by a set of arrows 1440pointing to the signal 1432. The highlighting indicates that blood flowhas been identified using the locating algorithm, which is typicallyassociated with a search mode, as previously described. The image of thevirtual surface 1416 in FIG. 14B can be simplified to show an outline ofthe virtual surface 1416 and the detected blood flow signal 1432, asillustrated in FIG. 14C. The image, while not displaying any of the lookdirections 210, focuses the operator on the detected blood flow signal1432. The image of FIG. 14C is provided after the acquisition/track modehas been entered. That is, after interrogating the region of interest bysurveying a plurality of sample locations (i.e., search), blood flow inthe search region has been identified (i.e., acquisition). A circle 1450within the blood flow signal 1432 is highlighted and represents theblood flow signal being tracked by the system. The image of FIG. 14C canbe switched to the image of FIG. 14D, which illustrates a PMD image 1460and a spectrogram image 1470 at the depth of interest for the blood flowsignal being tracked. As previously discussed, the PMD image 1460 can beused to display blood flow information in a time domain for the depth ofinterest.

As illustrated in the present example, the functionality of locating anacoustic window, automatically detecting blood flow, automaticallytracking blood flow, and further providing a PMD/spectrogram image ofthe blood flow can be combined to provide a tool for easily detectingand locating blood flow information, and monitoring the same forapplication in TCD.

From the foregoing it will be appreciated that, although specificembodiments of the invention have been described herein for purposes ofillustration, various modifications may be made without deviating fromthe spirit and scope of the invention. For example, the compound-modeDoppler virtual surface previously discussed can be constructed byelectronically steering an ultrasound beam to interrogate a plurality ofsample locations in a region of interest. In one modification, theentire compound-mode Doppler virtual surface can be swept or scannedover a larger region. That is, generally, a compound-mode Dopplervirtual surface is constructed while the probe is pointed and positionedat a first location. As the probe orientation or location is moved, thevirtual surface will be constructed for the region at which the probe isnow aimed. Thus, the entire virtual surface can be swept to cover alarger region of interest. Accordingly, the invention is not limitedexcept as by the appended claims.

1. A method for locating blood flow in a region of interest using aDoppler ultrasound system having a multi-element transducer, the methodcomprising: steering an ultrasound beam to deliver ultrasound to andreceive echo signals from a plurality of sample locations spatiallyarranged across the region of interest in two lateral dimensionsrelative to a reference beam axis extending from the multi-elementtransducer; for each sample location, processing the echo signals toextract Doppler shift signals and generate Doppler shift data therefromand accumulating the Doppler shift data for the plurality of samplelocations; determining from the Doppler shift data accumulated for theplurality of sample locations the presence of blood flow in the regionof interest; and identifying the location in the region of interest atwhich the presence of blood flow is detected.